Biological substance analysis methods based on optical means have risen in popularity in the last couple of decades. Common to all these methods is that chemical interactions between the bio-molecules produce changes that affect some measurable optical property, such as the emission spectrum, absorption spectrum or index of refraction. The changes in the optical properties can occur in the analyte itself or through a mediator such as the surface on which the interaction takes place. These changes are then monitored using a beam of incoming light (usually laser light) which in turn changes the outgoing light spectrum (e.g., fluorescence), intensity (e.g., absorption), or phase (e.g., surface plasmon resonance and any kind of interferometric method).
While most of these optical bio-analysis methods have found niche applications and markets, one method that became highly popular and influential was microarray optical fluorescence scanning Such optical scanning has enabled running tests on tens of thousands of miniature samples in a relatively short period of time. The major advantages of this method include: a) performance (sensitivity and signal to noise ratio (SNR)); b) speed; and c) miniaturization of the sampled analyte. These parameters define the efficiency and superiority of the method.
Currently microarray elements are spotted on top of a flat substrate chip usually made of glass, plastic or epoxy. Subsequently, the chip is scanned using confocal scanning systems where the exciting light and the resulting fluorescence light are both shined and collected from above and analyzed using a single photo-multiplier (PMT) detector. This arrangement suffers from several inherent limitations including a very short interaction length between the bio-sample and the light (usually a single mono-layer). This limits the signal strength and thus the SNR. Another limitation is a high background or noise due to the fact that the back-reflected light and the emitted fluorescent light travel in the same direction. A further limitation is high sensitivity to both the planarity and the position of the chip that need to be maintained in focus. Still another limitation is slow operation due to the need to have large enough number of ‘pixels’ (scanned spots) within every sample and long enough integration time. Yet another limitation is the need for a complicated optical and mechanical structure that entails bulky and expensive systems.
Another optical bio-analysis method is waveguide based bio-sensors. Bio-sensing based on waveguides has been around for a while. These biosensors can be divided into three main categories. The first involve slab waveguide fluorescence excitation with light collection from above or below the chip. In this arrangement the bio-analyzed spots are located on the surface of a chip that contains a single slab-waveguide. Light is coupled into the waveguide using a lens or a grating that excites the entire chip with all its bio-analyzed spots simultaneously. The fluorescence is collected using an optical imaging system and a charge-coupled device (CCD) detector from above or underneath the chip. One drawback of this kind of system is relatively poor performance due to uniformity of excitation as well as collection of the light. This leads to non-repeatable results. Another drawback is high noise levels due to crosstalk between the different spots. A further drawback is that large spots and relatively small numbers of elements are required to generate a signal large enough for efficient imaging with the CCD. Yet another drawback is the long integration time to overcome SNR issues. Examples of the above method are described in U.S. Pat. Nos. 5,814,565; 6,911,344 and 6,395,558.
A second waveguide based bio-sensor utilizes an interferometric optical device. In this case, channel waveguides are used together with interferometric devices such as Mach Zehnder interferometers (MZI) or ring-resonators. These sensitive interferometric devices sense the change in the index of refraction due to binding of the bio-molecules near a waveguide surface. The major problems associated with this type of system include non-specificity due to inability to recognize the exact reason for the index change, which may occur from deposition of other material as well as temperature changes. Another problem is a very slow speed in addressing the different elements which disqualifies this method for running large numbers of element arrays. Examples of the above method are described in U.S. Pat. Nos. 5,494,798 4,515,430, 5,623,561 and 6,618,536.
A third waveguide based bio-sensor utilizes surface plasmon resonance (SPR). Here, in one example, a thin gold layer is deposited on top of a glass substrate. The bio-analyzed sample on top of the gold induces changes in the refractive index above the gold layer, thus changing the resonant angle for generating surface plasmons along the gold layer. The plasmon generation is detected as an enhanced peak in the reflected beam. Examples of the SPR method are covered, for example, in U.S. Pat. No. 6,956,651 B2. Other types of optical bio-sensors and array scanners exist such as described in U.S. Pat. No. 6,396,995 B1.
One aspect common to all of these waveguide based sensors is the need to initially couple light into the waveguide. Since all of these optical waveguides have miniature cross-sections ranging from 100 micrometers down to a fraction of a micrometer, the coupling of light into the waveguide involves specialized optics for focusing the light, fine mechanical alignment for accurately placing the light source relative to the waveguide, and specialized glues to bond all components in place without interfering with the light. This process adds in most of these cases a considerable cost and complexity to the entire system.
In a large number of these optical waveguide applications, the light travels in the waveguide in the form of short pulses. These pulses can be as short as a pico (10−12) second and as long as a few milli (10−3) seconds. Moreover, these pulses can be all of the same wavelength or can be a combination of many different wavelengths. These pulses are generated by modulating one or more light sources which were initially coupled to the optical waveguide. If pulses at more than one wavelength are required, a combiner e.g., an Arrayed Waveguide Grating (AWG) must be added to the system.
In various applications, for example, biological analysis or detection systems, the optical waveguide may be part of a low-cost, consumable chip. Waveguide-based optical detection systems are disclosed, for example, in U.S. Pat. Nos. 7,951,583 B2, 8,187,866 B2 and 8,288,157 B2, each of which are hereby incorporated in their entirety by reference. In a system making use of such consumable chips, light needs to be coupled time after time to new chips. In such cases the cost and complexity of the light coupling technologies known in the art (see, for example, U.S. Pat. Nos. 4,881,789, 5,734,768, 5,600,744, 5,581,646, 5,444,805, 5,217,568, 5,121,457, 5,077,878, 4,744,623 and 4,478,485) are intolerable.
In addition, the optical signal generated on the sensing chip in all these applications is the result of one or more optical phenomena such as fluorescence, luminescence, absorption, scattering and optical phase change. Changes detected in the output light are indicative of the presence of the targets being detected. Nevertheless, the output light may also be affected by additional factors (e.g. temperature and aging) which may skew the results compromising the overall accuracy and precision of the detection system. Finding a method for normalizing the detected signal to overcome the effect of these uncontrolled factors is critical for maintaining the performance of the detection system.